Plga/hydroxyapatite composite bone grafts and method of making

ABSTRACT

The present invention involves tissue engineering constructs made from a new composite bone graft material made from biocompatible poly(D,L-lactic-co-glycolic acid) (PLGA) and bioceramic particles exposed on its surface using a gas foaming particle leaching (GF/PL) method and infused with collagen. Methods and apparatus for of forming scaffolds are also disclosed.

CROSS REFERENCE TO RELATED APPLICATION

This application is a Continuation application claiming priority to U.S.patent application Ser. No. 13/772,966, filed on Feb. 21, 2013 whichclaims priority to U.S. provisional patent application No. 61/601,281filed on Feb. 21, 2012 and cross references and incorporates byreference patent application Ser. No. 12/279,172, PCT/US2007/05693, Ser.No. 12/279,172 and 60/767,137. The contents of all references, patentsand patent applications are expressly incorporated by reference.

SUMMARY OF THE INVENTION

This invention is a novel biomaterial including formed shapes that areespecially useful in tissue engineering applications involving bone. Thecomposite comprises a polymeric scaffold, most preferablypoly(D,L-lactic-co-glycolic acid) (PLGA) and particles of a bioactivebioceramic such as hydroxyapatite (HA), triCalcium phosphate (TCP),calcium sulfate or bioglass or a combination thereof, wherein theceramic is highly exposed on the biomaterial surface. In one embodiment,a collagen solution is forced or infused through the biomaterial usingpressure, centrigution or a vacuum. A further embodiment of thisinvention involves a ceramic, such as apatite, that is fastly, highly,and uniformly coated on the biomaterial surface. This new biomaterial isadvantageous because it promotes bone cell propagation and ingrowthbetter than current materials. The biomechanical properties of thisbiocomposite in block form are far superior to hydroxyapatite blocks, oreven to the natural bone. The material when milled is also suitable foruse as a powder or granular material as a bone filler or cement, forfilling spaces from a few mm³ in volume to larger cm³. Volumes of boneneeding replacement can be measured using imaging technology and exactreplacements produced through computer-aided design (CAD) in conjunctionwith computer aided manufacturing (CAM).

While PLGA scaffolds have been used in the past, such scaffolds were notoptimal for the present applications because the porosity of thescaffolds made them degrade too quickly. HA blocks are widely available,but they are brittle and often break when anchored to the host bone witha screw. Following the methods and specifications described herein, onecan fabricate scaffolds from PLGA which will be resident in the bodylong enough for the desired biological growth to occur before the

BACKGROUND

The ideal bone graft would replace bone defects, such as those fromdisease or trauma, with a material that allows bone cells to grow intothe affected area, thus restoring the bone to its original condition.Currently, autografts are the best material for bone repair because theyare biocompatible and there is little risk of disease transfer. However,the downside of autografts is that a separate operation must beperformed to remove the person's own bone. Allografts, which consist ofbone from another person/cadaver, as well as xenograft, (bone fromanother animal species) are also available but carry the risk of immuneresponse and disease transfer that could lead to ultimate failure.

In order to solve the problems associated with bone grafts, manyresearchers have tried to develop artificial substances for bone grafts.These artificial biomaterials need to possess several qualities in orderto be successful. First, the material must be porous to allow room fornew bone to grow into the implant site. Second, it must maintainmechanical strength similar to native bone. Finally, the artificialbiomaterial needs to be osteoconductive; that is, it must allow bonecells to attach and propagate on its surface, as it resorbs.

Some of the materials that have shown promise as bone grafts includecalcium phosphate ceramics such as hydroxyapatite and tricalciumphosphate. These particular ceramics are quite biocompatible becausethey have characteristics similar to native bone mineral. However, theyare hard to shape and do not possess the same mechanical properties asbone. They are quite brittle and require extremely delicate handlingwhen shaping or drilling to avoid breaking the material. Hydroxyapatitedegrades very slowly, which inhibits new bone from forming.

Another type of material that has sparked some interest is the use ofdegradable polymer. Polymers easy to shape and degrade at a predictablerate, thereby allowing new bone growth to replace it. Some examples ofdegradable polymers are poly(glycolic acid), poly(L-lactic acid), andpoly(D,L-lactic-co-glycolic acid). Although they are easily formed andhave good mechanical strength, degradable polymers alone are not idealfor bone grafts because they are not very osteoconductive. New bone willnot attach well or grow well into this material.

Synthetic polymers which can be used in the present invention includepoly(hydroxy acids) such as poly(lactic acid) (PLA), poly(L-lactic acid)(PLLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid)(PLGA), poly(L-lactic acid-co-glycolic acid) (PLLGA), polyanhydrides,polyorthoesters, poly(ester amides), polyamides, poly(esterethers)polycarbonates, polyalkylenes such as polyethylene andpolypropylene, polyalkylene glycols such as poly(ethylene glycol) (PEG),polyalkylene oxides such as poly(ethylene oxide) (PEO), polyalkyleneterephthalates such as poly(ethylene terephthalate), polyvinyl alcohols(PVA), polyvinyl ethers, polyvinyl esters such as poly(vinyl acetate),polyvinyl halides such as poly(vinyl chloride) (PVC),polyvinylpyrrolidone, polysiloxanes, polystyrene (PS), polyurethanes,derivatized celluloses such as alkyl celluloses, hydroxyalkylcelluloses, cellulose ethers, cellulose esters, nitro celluloses,polymers of acrylic acids, such as poly(methyl(meth)acrylate) (μmMA),poly(ethyl(meth)acrylate), poly(butyl(meth)acrylate),poly(isobutyl(meth)acrylate), poly(hexyl(meth)acrylate),poly(isodecyl(meth)acrylate), poly(lauryl(meth)acrylate),poly(phenyl(meth)acrylate), poly(methyl acrylate), poly(isopropylacrylate), poly(isobutyl acrylate), poly(octadecyl acrylate) (jointlyreferred to herein as “polyacrylic acids”), and copolymers and mixturesthereof, polydioxanone and its copolymers, polyhydroxyalkanoates,polypropylene fumarate), polyoxymethylene, and poloxamers.

The polymers can optionally include one or more photopolymerizablegroups. The polymers can also be derivativatized. For example, thepolymers can have substitutions such as alkyl groups, alkylene groups,or other chemical groups. The polymers can also be hydroxylatedoxidized, or modified in some other way familiar to those skilled in theart. Blends and co-polymers of these polymers can also be used.

Preferred non-biodegradable polymers include ethylene vinyl acetate,polyacrylic acids, polyamides, and copolymers and blends thereof.

Preferred biodegradable polymers include poly(hydroxy acids) such asPoly lactic acid (PLA), poly glycolic acid (PGA), Poly lacticco-glycolic acid (PLGA), and copolymers with polyethylene glycol (PEG);polyanhydrides, poly(ortho)esters, polyurethanes, poly(butyric acid),poly(valeric acid), poly(lactide-co-caprolactone), trimethylenecarbonate, and the polymers described in Hubbell et al., U.S. Pat. Nos.5,654,381; 5,627,233; 5,628,863; 5,567,440; and 5,567,435. In general,these materials degrade in vivo by both non-enzymatic and enzymatichydrolysis, and by surface or bulk erosion.

Preferred water-soluble polymers include polyethylene oxides,polyethylene glycols, ethylene oxide-propylene oxide copolymers(poloxamers and poloxamines), polyvinyl alcohols, polyvinylpyrrolidones,poly(acrylic acids), and copolymers and blends thereof.

Natural polymers that can be used in the invention includepolysaccharides such as alginate, dextran, and celluloses; collagens,including derivatized collagens (e.g., alkylated, hydroxylated,oxidized, or PEG-lated collagens, as well as collagens modified by otheralterations routinely made by those skilled in the art); hydrophilicproteins such as albumin; hydrophobic proteins such as protamines, andcopolymers and mixtures thereof. In general, these materials degrade byenzymatic hydrolysis, by exposure to water in vivo, or by surface orbulk erosion.

Preferred bioadhesive polymers include polyanhydrides and polyacrylicacids. In one embodiment, reactive groups on the polymers, for example,hydroxy, amine, carboxylic acid, thiol, anhydride, ester and vinylgroups are reacted with reactive groups on agents to be incorporatedinto the polymer matrix. For example, bioactive compounds such asproteins contain reactive amine groups which can be coupled withreactive carboxylic acid, ester, or anhydride groups on the polymer toform polymers that are covalently bonded to the compounds. In anotherembodiment, ion pairs are formed between acidic or basic groups on apolymer and basic or acidic groups on a bioactive compound to form apolymer that is ionically bonded to the compounds. Those of skill in theart can readily determine an appropriate bioactive compound and polymerto couple by forming ionic or covalent bonds, and can also readilydetermine appropriate reaction conditions for forming such bonds.

One factor to be considered when selecting an appropriate polymer is thetime required for in vivo stability, i.e., the time in which the polymermatrix is required to degrade, in those embodiments in which the matrixis used in vivo. Preferably, the polymer matrix exhibits an in vivostability between approximately a few minutes and one year. When usedfor drug delivery, the in vivo stability is preferably between a fewhours and two months. When used for tissue engineering, the in vivostability is preferably between one week and several months.

The art has used blocks of hydroxyapatite tri calcium phosphate (HA TCP)as a bone graft material. Such materials are extremely brittle andfracture when drilled or a screw is inserted. Additionally, HA TCP whenused alone in blocks has limited porosity, if any, and tends to getencapsulated as a foreign body when implanted. As such, HA TCP Block isnever truly integrated into the existing bone except at a very narrowmargin at the surface of the block. As a result, it does not gain thestrength of a graft fabricated from harvested bone.

It is possible to make a composite using a phosphate ceramic inconjunction with a degradable polymer. Small particles of ceramic can beincluded within the polymer scaffold material. These particles will bepartially exposed on the surface of the biomaterial, thereby making thematerial more osteoconductive.

Most related methods for making a polymer/ceramic scaffold biomaterialuse organic solvents. This can be highly disadvantageous because someresidual solvent may remain in the material. Almost all organic solventsare detrimental to cell and tissue growth. Also, it has been noted thatthese processes may actually leave behind a thin film of polymer thatcoats the ceramic particles that are supposed to be exposed on thesurface. This unintentional thin film disrupts the osteoconductivenature of the ceramic portion of these biomaterials.

Shaping polymer base scaffolds has presented significant challengesbecause the use of mechanical cutting and shaping devices such as drillsor saws melts the polymer distorting the surface. In particular, theexposed hydroxyapatite is occluded rendering the material a lesseffective bone replacement. Conventional abrasives such as aluminumoxide or carborundum generally cannot be used because they willcontaminate the scaffold.

The invention disclosed herein addresses the problems by describing apolymer/ceramic biomaterial comprised of degradable polymer and ceramicwherein the ceramic is highly exposed on the surface of the biomaterialand the biomaterial is fabricated with no use of organic solvents. Thematerials are infused with collagen, providing further attachment pointfor osteogenic cells. Furthermore, an additional layer of a mineral,such as apatite, can be coated on the surface of the biomaterial in anadherent, fast, uniform fashion. Finally, granules of thepolymer/ceramic biomaterial with additional ceramic coating can befabricated.

All references cited within this application are expressly incorporatedby reference in their entirety.

BRIEF SUMMARY OF THE INVENTION

A preferred embodiment of the present invention is a biomaterialcomprised of poly(D,L-lactic-co-glycolic acid) (PLGA), hydroxyapatite,and a possible coating of apatite. It is suitable as an artificial bonegraft material. The said biomaterial is formed using a gas foamingmethod. GF introduces gas bubbles into the polymer matrix by saturatingthe polymer with gas at high pressure, and then reducing the pressureback to ambient conditions at a sufficiently fast rate to induce bubbleformation. In a preferred embodiment CO2 gas is used, but any gas whichis non-reactive with the biomaterial or its components may be used. Inthe present invention, hydroxyapatite serves as a porogen. Optionally,an additional porogen can be used to form pores, but such additionalporogens have been found to create too much surface area in the PLGAleading to rapid degradation. Such additional porogens can be anybiologically acceptable salt, gelatin, saccharose crystals or any othersolid agent which can be solubilized in a solution which does notdegrade PLGA or hydroxyapatite. While NaCl is a preferred salt, the saltparticles can be any salt that can form crystals or particles having adiameter from 0.1 mm to 2 mm, which is easily removed from and does notreact with the polymer, and is non-toxic if some residue remains in thepolymer after leaching. Suitable salts include sodium salts, such assodium chloride, sodium tartrate and sodium citrate, and other watersoluble salts or compounds which remain insoluble in the polymer.Alternatively other particle forming agents which are capable of beingwashed out of the biomaterial may be used. Such agents can includeproteins such as gelatin and agarose, starches, polysaccharides such asalginate and other polymers. In the preferred embodiment, pore formingparticles are water soluble and leached out of the matrix usingdistilled water. Both gas foaming and particle leaching leave behindvoids, which form the pores of this biomaterial matrix.

A preferred embodiment is made by combining particles of polymer andbioceramic, in certain ratios and then using the GF method. The size andamount of each particle will determine the general and interconnectedporosity of the final biomaterial. Initially the polymer and bioceramicparticles are sieved to obtain particles with a specific size. Thenthese particles are combined in certain ratios and loaded into a mold.The mixture is then exposed to high pressure CO2 gas (around 2000 psi)for 3-4 hours at a temperature of around 200 degrees F. It is believedthat pressures as low as 400 psi and as high as 4000 psi will also work.During this time, the CO2 saturates the scaffold After 3-4 hours, theCO2 gas pressure is decreased to ambient pressure at a rate fast enoughto induce nucleation and growth of CO2 bubbles, which form pores withinthe polymer scaffold portions of the biomaterial. If present, theoptional sodium chloride particles are subsequently leached out of thematerial by immersing the scaffold in distilled water for a sufficientamount of time to dissolve the salt, thus leaving voids formerlyoccupied by the sodium chloride particles. While failing to remove thepore forming agent (salt) is not detrimental, more than one immersion inclean distilled water may be used to ensure removal of the pore formingparticles. The final material is highly porous with bioceramic particlesexposed on the surface of its polymer network.

The compositions are the present invention are collagen infused byplacing them in a suitable vessel containing a solution of collagen andforcing the collagen into the pores via pressure, centrifugation or useof a vacuum. (negative pressure).

Furthermore, coating the surface of the polymer/bioceramic scaffold witha bone-like apatite using a biomimetic process can increase itsosteogenic potential. This biomimetic process involves soaking thebiomaterial in a solution of simulated body fluid (SBF) that hasappropriate concentrations of ions dissolved in solution. Certain ionswill precipitate on the surface of the biomaterial and form an apatitemineral coating.

In another embodiment, the polymer/bioceramic biomaterial may be groundup and sieved to collect granules with a certain size. These granulesmay then be soaked in the SBF and receive the apatite coating thatenhances its osteogenic properties.

A bone graft material according to the present teachings can be providedin the form of a bone paste, a shaped solid, or a dry pre-mix useful forforming such a paste or solid. The phrase “bone paste” refers to aslurry or semi-solid composition of any consistency that hardens to forma solid structure, and thus includes, e.g., bone plasters, putties,adhesives, cements, bone void fillers, and bone substitutes. As aresult, the bone paste can be any composition capable of being injected,molded, painted, suffused, or placed into contact with a bone surface invivo. The “shaped solid” can take any form, including a pellet that canbe placed into a bone void or into contact with a bone surface in vivo.The dry pre-mix can be provided in the form of a powdered and/orgranular material.

In another embodiment, the biomaterial is further infused with nanocollagen fibers by forcing a collagen solution through the biomaterial.

In another embodiment the biomaterial is shaped using CAD/CAMtechnologies to more precisely replicate a region of bone requiringreplacement.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a black and white photograph of a bone graft of the presentinvention.

FIG. 2 is a drawing of a mold having a Teflon sheath inserted.

FIG. 3 is a top view of the mold showing Teflon sheath.

FIG. 4 is a drawing of the mold showing Teflon sheath removed.

FIG. 5 is a top view of the assembled mold.

FIG. 6 is a side view of the reaction vessel.

FIG. 7 is a side view of the reaction vessel with the secondarycontainment installed.

FIG. 8 is a drawing of the top of the reaction vessel cap and thepressure pump.

FIG. 9 is a drawing of the pressure pump.

FIG. 10 is a drawing of the reaction system including the CO2 tank.

FIG. 11 is a bottom view of the reactor cap.

FIG. 12 is a view of the control panel.

FIG. 13 is a CO2 phase diagram.

FIG. 14 is a close black and white photograph of a powder form of thepresent invention.

FIG. 15 is a close black and white photograph of the block form of thepresent invention.

DETAILED DESCRIPTION OF THE INVENTION

The present invention is a novel biomaterial with specialcharacteristics that allow it to perform well as a bone graft material.It is comprised of a polymer scaffold, preferably a degradable poly(D,Llactic-co-glycolic acid) polymer with incorporated bioceramic particlesand made by a gas foaming (GF) method. A further embodiment of thisinvention describes the same biomaterial infused with collagen and/orwith an adherent, highly uniform apatite coating.

The method of constructing a PLGA polymer scaffold using GF/PL isdescribed thoroughly in the journal article titled, “Open porebiodegradable matrices formed with gas foaming” (Harris L D, Kim B S,and Mooney D J; J Biomed Mater Res, 42, 396-402, 1998). This entirearticle is hereby incorporated by reference. The research reported inthis article found that the porosity and pore size of the PLGA scaffoldcan be controlled by the salt/PLGA ratio and respective particle sizes.Also, the pores of the matrix are interconnected and highly uniform. Inthis manner, a useful scaffold can be created without the use of organicsolvents or high temperatures. The present inventor has discovered thatthe particle leaching of Kim is not required to produce suitable poresand in fact products scaffolds that degrade too quickly due to excessiveporosity.

While methods for constructing a polymer scaffold using GF are known,the use polymers with bioceramic particles as porogens in connectionwith high pressure gas foaming has not been taught in the prior art. KimS S, Park M S, Jeon O, Choi C Y, and Kim B S;Poly(lactide-co-glycolide)/hydroxyapatite composite scaffolds for bonetissue engineering, Biomaterial, 27, 1399-1409, available online Oct. 5,2005 describes the addition of nano hydroxyapatite particles to a PLGAscaffold. This article is also hereby incorporated by reference.

The most significant modification of the prior art methods is theomission of porogens which are washed out after scaffold formation andthe infusion of collagen throughout the pores.

L-Lactide/caprolactone copolymer/HATCP composites were prepared with75:25 Lactide/caprolactone copolymer particles (diameter=100-200microns, molecular weight=100,000 Purac Biomaterials), HA TCP particles(diameter=approximately 100−1000 nm, 40% HA to 60% TCP). The polymerparticles were mixed with the HA TCP particles. The PLGA/HATCP massratio ranged from 80:20 to 50:50 PLGA to HA by weight. The mixture wasloaded into a mold and exposed to high pressure CO2 gas (2000 psi) for3-4 hours to saturate the polymer with the gas. Temperature in thereaction vessel is maintained at around 200 degrees F. to maintain CO2in the supercritical range. Then, decreasing the gas pressure to ambientpressure created a thermodynamic instability which led to the nucleationand growth of CO2 pores within the polymer scaffolds.

In a preferred embodiment the pressure is supplied by a reaction vessel25. Referring to FIGS. 2-11, the reaction vessel 25 is preferablyconstructed out of stainless steel or other suitable material capable ofwithstanding the pressures recited herein. The vessel consists of a body41 which defines a chamber 39. The chamber 39 is closed and sealed atone end and open on another end. In the figures the chamber is sealedwith a bottom cap 28, which is bolted to a bottom flange 27 which ispart of body 41. The open end is sealed by bolting a cap 31 whichattached to a top flange 29 which is part of body 41. Gaskets or O-ringsare used as required for sealing. Referring to FIG. 11 the caps and/orthe flanges can include grooves 44 for one or more O-rings 45. While thephotos show bolts used to hold the caps to the reaction vessel, anyother means capable of handling the pressure may be used such as clamps,thumbscrews, etc.

A means is provided for increasing the pressure in the chamber 39. Suchmeans can be externally mounted pumps 33 or pumps which are integratedinto the reaction vessel. The vessel contains an inlet 34 which isconnected to a manifold to control pressure in the chamber.

In the present embodiment, the pump 33 is connected via hose 35 to CO2tank 41. The pump can be any pump suitable for pressurizing fluid or gasto the pressures recited herein. In the present embodiment the pump is acustom made pressure pump which is capable of generating pressures ofbetween 2500-3000 psi. Such pumps are known in the art. The systemincludes a pressure gauge 32 for measuring pressure in chamber 39 in thereaction vessel 25. The pressure gauge 32 can be mounted anywheresuitable for measuring the pressure including the caps, body or inputlines. In these Figures the pressure gauge 32 is mounted to the top cap.

Referring to FIGS. 6,7 and 10, a secondary containment case 26 isprovided around the reaction vessel 25 to hold the heating device.

The secondary containment case 26 also houses a heater 42 which is usedto elevate the temperature in the reaction vessel 25. The heater is anelectric resistance heating element capable of maintaining temperaturesin the reaction vessel of around 200 degrees F. It is contemplated thatother sources of heat could be applied in an industrial setting such asinfrared, steam or heated water.

While the above materials are preferred, it is believed that the presentinvention will work with polymers of diameters from 50-500 microns andwith bioceramic particles having diameters of 50-500 microns and saltparticles having diameters from 50-300 microns. It is believed that theratios of polymer to hydroxy apatite to NaCl can vary by as much as 50%without deviating from the spirit of this invention.

The process for creating these polymer/bioceramic composite biomaterialscan be summarized by the steps of: (1) obtaining polymer in theappropriate particle size range (grinding small particles if necessary),(2) optionally sieving the polymer and bioceramic particles to yieldparticles with a 100-200 microns diameter, (3) mixing the particles ofpolymer and bioceramic in a mass ratio from about 4:1 polymer tobioceramic to about 1:1 polymer to bioceramic, (4) loading the mixtureof particles into a mold, (5) compressing the mixture with a very highpressure CO2 gas long enough to saturate the disk, (6) decreasing thepressure on the disk until it returns to ambient pressure.

The rate at which pressure is vented from the reaction vessel determinesthe pore size. The faster the pressure release, the larger the poresize. It has been determined that decreasing the pressure to ambientover a period of 15-25 minutes yields desirable pore sizes. The fasterpressure is released the larger the pore sizes. One of skill in the artwill be able to determine the rate of release based on the specificcomposition of the scaffolds being produced.

Optionally, a further embodiment of the invention involves forming auniform mineral coating of apatite on the surface of thepolymer/bioceramic biomaterial. This apatite layer enhances theosteogenic potential of the biomaterial scaffold.

The apatite layer is created by incubating the bone graft in an ion richsimulated body fluid (SBF) solution. The solution is prepared bydissolving reagent grade NaCl, NaHCO3, Na2SO4, KCl, K2HPO4, MgCl2.6H2O,and CaCl2.2H2O in distilled deionized water. 1×SBF has the same ionconcentrations as blood plasma while 5×SBF has ion concentrations fivetimes greater than blood plasma. The pH is adjusted to 6.4 withtris(hydroxymethyl) aminomethane.

The described PLGA/hydroxyapatite biomaterial can be coated with apatiterelatively quickly because the exposed hydroxyapatite particles act asnucleation sites for the growth of the mineral apatite layer in SBFsolution.

A further embodiment of this invention involves the coating ofPLGA/nanohydroxyapatite particles (rather than scaffolds) with abiomimetic, adherent, and uniform apatite coating. These particles willthen be soaked in SBF solution to coat them with a uniform layer ofbiomimetic apatite.

Most of the previous methods for fabricating polymer/bioceramiccomposite scaffolds, such as the solvent casting and particulateleaching (SC/PL) method or the phase separation method, use organicsolvents. However, residual solvents in the scaffolds may be harmful totransplanted cells or host tissues. Furthermore, the polymer coating onthe ceramics created by polymer solutions may hinder the exposure of theceramics to the scaffold surfaces, which could decrease the chance thatosteogenic cells make contact with the bioactive ceramics.

The preferred embodiment of the present invention relies on gas forming(GF) methods to fabricate polymer bioceramic composite scaffolds forbone tissue engineering. This method efficiently exposes the bioceramicon the scaffold surfaces and avoids the use of organic solvents.Bioceramic particles used to fabricate the composite scaffolds areapproximately 100 nanometers to 1000 microns in size.

Example 1 Process

L-Lactide/caprolactone copolymer/HA TCP composites were prepared with75:25 Lactide/caprolactone copolymer particles (diameter=100-200microns, molecular weight=100,000 Da, Purac Biomataerials), HA TCPparticles (diameter=approximately 100−1000 nm, 40% HA to 60% TCP). Thepolymer particles were mixed with the HA TCP particles. The polymer/HATCP mass ratio ranged from 80:20 to 50:50 polymer to HA TCP by weight.The mixture was loaded into a mold and exposed to high pressure CO2 gas(2000 psi) for 3-4 hours to saturate the polymer with the gas.Temperature in the reaction vessel is maintained at around 200 degreesF. to maintain CO2 in the supercritical range. Then, decreasing the gaspressure to ambient pressure created a thermodynamic instability whichled to the nucleation and growth of CO2 pores within the polymerscaffolds.

The specific ratios of polymers is not critical, but plays a role inabsorption. 50:50 ratios of co polymers tend to absorb faster. Changingthe ratios from 50:50 can delay absorption by as much as 10%. Theinventor has determined that

Temperature in the reaction vessel is maintained at around 200 degreesF. to maintain CO2 in the supercritical range. Then, decreasing the gaspressure to ambient pressure created a thermodynamic instability. Thisled to the nucleation and growth of CO2 pores within the polymerscaffolds.

One of skill in the art will understand that temperature and pressurecan be adjusted to yield a suitable product. Pressures can range from400 to 4000 psi. The temperatures within the reaction vessel shouldequal or exceed the glass transition point for the polymers being usedbut should be below the temperature at which undesirable degradation ofany materials in the reaction vessel occurs. It is preferred that thetemperature and pressure be maintained such that the CO2 achieves asupercritical state. When in a supercritical state, substances are at atemperature and pressure above the thermodynamic critical point and havethe ability to diffuse through solids like a gas, yet retain the abilityto solubilize materials like a liquid. A representative phase diagramfor CO2 is provided as FIG. 13.

In a preferred embodiment the pressure is supplied by a reaction vessel25. Referring to FIGS. 2-11, the reaction vessel 25 is preferablyconstructed out of stainless steel or other suitable material capable ofwithstanding the pressures recited herein. The vessel consists of a body41 which defines a chamber 39. The chamber 39 is closed and sealed atone end and open on another end. In the figures the chamber is sealedwith a bottom cap 28 which is bolted to a bottom flange 27 which is partof body 41. The open end is sealed by bolting a cap 31 which attached toa top flange 29 which is part of body 41. Gaskets or O-rings are used asrequired for sealing. Referring to FIG. 10, the caps and/or the flangescan include one or more grooves for O-rings or seals. In a preferredembodiment, the seal is Teflon. While the photos show bolts used to holdthe caps to the reaction vessel, any other means capable of handling thepressure may be used such as clamps.

A means is provided for increasing the pressure in the chamber 39. Suchmeans can be externally mounted pumps 33 or pumps which are integratedinto the reaction vessel. The vessel contains an inlet 34 which isconnected to a manifold to control pressure in the chamber.

In the present embodiment, the pump 33 is connected via hose 35 to CO2tank 41. The pump can be any pump suitable for pressurizing fluid or gasto the pressures recited herein. In the present embodiment the pump is acustom made pressure pump which is capable of generating pressures ofbetween 2500-3000 psi. Such pumps are known in the art. The systemincludes a pressure gauge 32 for measuring pressure in chamber 39 in thereaction vessel 25. The pressure gauge 32 can be mounted anywheresuitable for measuring the pressure including the caps, body or inputlines. In these Figures the pressure gauge 32 is mounted to the top cap.

Referring to FIGS. 6, 7 and 10, a secondary containment case 26 isprovided around the reaction vessel 25. Such secondary containment isprovided for added safety in the event of a rupture in the body of thereaction vessel. The secondary containment case can be fabricated out ofany suitable material which can contain shrapnel from the reactionvessel in the event of a breach. Any suitable ballistics material can beused including, ballistics blankets, suitable metals such as stainlesssteel, aluminum, steel or alloys, composites such as fiberglass, carbonfiber or Kevlar, or polymers. In the present embodiment, the containmentcase 26 is made out of Teflon.

The secondary containment case 26 also houses a heater 42 which is usedto elevate the temperature in the reaction vessel 25. The heater is anelectric resistance heating element capable of maintaining temperaturesin the reaction vessel of around 200 degrees F. It is contemplated thatother sources of heat could be applied in an industrial setting such asinfrared, steam or heated water.

While the above materials are preferred, it is believed that the presentinvention will work with polymers of diameters from 50-400 microns andwith bioceramic particles having diameters of 150-1000 microns andoptional salt particles having diameters from 50-300 microns.

The process for creating these polymer/bioceramic composite biomaterialscan be summarized by the steps of: (1) obtaining polymer in theappropriate particle size range (grinding small particles if necessary),(2) optionally sieving the polymer and bioceramic particles to yieldparticles with a 100-200 microns diameter, (3) mixing the particles ofpolymer and bioceramic in a mass ratio from about 4:1 polymer tobioceramic to about 1:1 polymer to bioceramic, (4) loading the mixtureof particles into a mold, (5) compressing the mixture with a very highpressure CO2 gas long enough to saturate the disk, (6) decreasing thepressure on the disk until it returns to ambient pressure.

The rate at which pressure is vented from the reaction vessel determinesthe pore size. The faster the pressure release, the larger the poresize. It has been determined that decreasing the pressure to ambientover a period of 15-25 minutes yields desirable pore sizes.

Optionally, a further embodiment of the invention involves forming auniform mineral coating of apatite on the surface of thePLGA/hydroxyapatite biomaterial. This apatite layer enhances theosteogenic potential of the biomaterial scaffold.

The apatite layer is created by incubating the bone graft in an ion richsimulated body fluid (SBF) solution. The solution is prepared bydissolving reagent grade NaCl, NaHCO3, Na2SO4, KCl, K2HPO4, MgCl2.6H2O,and CaCl2.2H2O in distilled deionized water. 1×SBF has the same ionconcentrations as blood plasma while 5×SBF has ion concentrations fivetimes greater than blood plasma. The pH is adjusted to 6.4 withtris(hydroxymethyl) aminomethane.

The described PLGA/hydroxyapatite biomaterial can be coated with apatiterelatively quickly because the exposed hydroxyapatite particles act asnucleation sites for the growth of the mineral apatite layer in SBFsolution. Although the method for coating of polymeric biomaterial withapatite by incubating the biomaterial in SBF solution is already known,accelerated coating by incubating polymeric biomaterial withnano-hydroxyapatites exposed on the biomaterial surface has not beentaught in the prior art.

Finally, the scaffolds were air-dried and then vacuum dried. The massincrease from apatite formation would be expressed as a percent increasecompared to the scaffold mass when incubated in a tris-buffer at thesame pH value, at the same temperature, and for the same time intervals.

The biomimetic apatite coating process is enhanced by introducingnano-sized hydroxyapatite nucleation sites and by using concentrated SBFsolution. This coating is advantageous because it conveys betterosteogenic properties to the polymer/bioceramic biomaterial.

A further embodiment of this invention involves the coating thescaffolds or powder with an adherent, and uniform apatite coating. Theparticles may be the product of a reaction process or be ground downfrom bulk polymer/bioceramic composite to a size of 30-2000 microns. Theparticles will be sieved to isolate particles with a more narrow sizedistribution depending on the desired application. These particles willthen be soaked in SBF solution to coat them with a uniform layer ofbiomimetic apatite.

Most of the previous methods for fabricating polymer/bioceramiccomposite scaffolds, such as the solvent casting and particulateleaching (SC/PL) method or the phase separation method, use organicsolvents. However, residual solvents in the scaffolds may be harmful totransplanted cells or host tissues. Furthermore, the polymer coating onthe ceramics created by polymer solutions may hinder the exposure of theceramics to the scaffold surfaces, which could decrease the chance thatosteogenic cells make contact with the bioactive ceramics.

The preferred embodiment of the present invention relies on gas forming(GF) methods to fabricate polymer/bioceramic composite scaffolds forbone tissue engineering. This method efficiently exposes the bioceramicon the scaffold surfaces and avoids the use of organic solvents. Toreduce the amount of bioceramic (HA degrades extremely slowly in vivo)required, and to increase the bioceramic exposure to the scaffoldsurface, bioceramic particles approximately 100 nm to 2 mm in sizerather than micro-sized particles, are used to fabricate the compositescaffolds. The most preferred range of bioceramic particles is betweenabout 150 microns to about 400 microns. The size of the particlesdepends in part on the rate of in vivo absorption. HA is known to absorbvery slowly, whereas TCP and calcium sulfate are known to be absorbedfaster. If HA is used alone, smaller particles are preferred, but whenfaster absorbing materials are used, larger particles are preferred.

The porosity of fabricated scaffolds can be measured using mercuryintrusion porosimetry (Autopore IV 9500, Micromeritics InstrumentCorporation, Norcross, Ga.). A contact angle of 1301 for mercury on thescaffold was used for this analysis. The pore structures of thescaffolds were examined using a scanning electron microscope (SEM, JEOL,Tokyo, Japan). Compression and tensile tests were performed with anInstron mechanical tester (Instron 4201, Instrons, Canton, Mass.). Thescaffold samples are cut into 1×1 cm2 for compression testing. Fortensile testing, the samples (1×1 cm2) can be attached to cardboardusing epoxy glue. The sample can be centered in a 7 mm slot in thecenter of the cardboard and then glued to standardize the gauge length.Compression and tensile tests can be performed with a constant strainrate of 1 mm/min. The moduli can be determined from the slopes in theinitial elastic portion of the stress-strain diagram. To examine thedistribution and extent of surface exposure of HA in the scaffolds, theHA exposed to the scaffold surface can be visualized with a hydrophilicdye (trypan blue, Sigma) staining. The residual dye can be removed bysonication in 100% ethanol. Afterwards, the surface of the PLGA/HAscaffolds can be examined with a microscope (Camscope, Samtech, Seoul,Korea). To examine the chemical composition of the scaffold surface, onecan carry out X-ray photoelectron spectroscopic (XPS; Sigma Probe,ThermoVG Scientific, West Sussex, UK) analyses, evaluating the O 1 s, C1 s, Ca 2p, and P 2p peaks. The residual pressure in the spectrometershould be 1.1×10⁻8 Pa, and an Mg anode (1.25 keV) powered at 250 W canbe used as an X-ray source. The constant pass energy was 23 eV. All XPSdata can be acquired at a nominal photoelectron takeoff angle of 551.The area of the XPS peaks can be determined after backgroundsubtraction, and the atomic percentage can be determined by normalizingthe peak area of each element by the total peak areas of all elements.

Osteoblasts can be isolated from the calvaria of neonatal (less than oneday old) Sprague-Dawley rats (SLC, Tokyo, Japan) by an enzymaticdigestive process. The calvaria can be isolated, and all connectivetissues were carefully removed. The parietal bones can be minced intopieces measuring about 1×1 mm² using sterile surgical scissors.Osteoblasts can be isolated by an enzyme solution containing 1.37 mg/mlcollagenase type I (Sigma) and 0.5 mg/ml trypsin (Sigma). Following 30min of incubation, the released cells are discarded to preventcontamination with other cell types. The minced bones are redigestedwith the enzyme solution for 30 min, and the supernatant is transferredto the culture medium, Dulbecco's Modified Eagles Medium (DMEM, GibcoBRL, Gaithersburg, Md.) containing 10% (v/v) fetal bovine serum (GibcoBRL), 1% (v/v) penicillin-streptomycin (Gibco BRL), 10 mMb-glycerophosphate (Sigma), 50 mg/ml L-ascorbic acid (Sigma), and 100 nMdexamethasone (Sigma). This process should be repeated three times, andthen finally the collected solution is centrifuged for 10 min at 1500rpm. Cells are plated into tissue culture flasks and cultured in ahumidified incubator at 37° C. with 5% (v/v) CO2.

The fabricated scaffolds are sterilized by ethylene oxide gas andpre-wetted in the culture medium for 12 h. Aliquots of 50 ml of the cellsuspension (4.0×10⁷ cells/ml, 2.0×10⁶ cells/scaffold) are seeded ontothe tops of the pre-wetted scaffolds. The scaffolds are left undisturbedin an incubator for 3 h to allow the cells to attach to the scaffolds.An additional 1 and 10 ml of culture medium are added to each scaffoldat 6 and 8 h, respectively. The cell/scaffold constructs are cultured ina humidified incubator at 37° C. with 5% (v/v) CO2 for eight weeks. Themedium was changed every day. Analytical assays are performed at 7, 14,28, and 56 days.

To determine the seeding efficiency and cell growth on the scaffolds,cell numbers are determined by quantitative DNA assays (n=3). DNA wasisolated using a Wizard Genomic DNA Purification kit (Promega, Madison,Wis.). For DNA isolation, the cell/scaffold constructs are washed twicewith phosphate-buffered saline. The specimens were placed in a 1.5-mltube and crushed with a homogenizer (PowerGen 125, Fisher Scientific,Germany). DNA is isolated according to the kit protocol, and DNA contentis measured with an ultraviolet absorbance spectrophotometer (JASCOV-530, Tokyo, Japan) at 260 nm. The cell numbers are calculated from aDNA standard curve of identical cells.

The alkaline phosphatase (ALP) production of osteoblasts cultured onscaffolds can be measured spectroscopically (n=3) using the methods ofEkholm M, Hietanen J, Tulamo R M, Muhonen J, Lindqvist C, Kellomaki M,et al. Tissue reactions of subcutaneously implanted mixture ofepsilon-caprolactone-lactide copolymer and tricalcium phosphate. Anelectron microscopic evaluation in sheep. J Mater Sci Mater Med 2003;14:913-8. The osteoblast/scaffold constructs are washed with PBS,homogenized with 1 ml Tris buffer (1 M, pH 8.0, Sigma), and sonicatedfor 4 min on ice. Aliquots of 20 ml are incubated with 1 ml of ap-nitrophenyl phosphate solution (16 mM, Sigma) at 30 1C for up to 5min. The production of p-nitrophenol in the presence of ALP is measuredby monitoring light absorbance at 405 nm.

The amount of calcium deposited in the cell-scaffold constructs can bemeasured using a previously reported method (n=3) of Jaiswal N,Haynesworth S E, Caplan A I, Bruder S P. Osteogenic differentiation ofpurified, culture-expanded human mesenchymal stem cells in vitro. J CellBiochem 1997; 64:295-312. After the cell-scaffold constructs are rinsedtwice with PBS and homogenized with 0.6 N HCl, calcium is extracted byshaking for 4 h at 4° C. The lysate was then centrifuged at 1000 g for 5min, and the supernatant is used to determine calcium content. Tomeasure the amount of calcium produced by the seeded osteoblasts, thecalcium content of the PLGA/HA scaffold itself is also measured, and thecalcium content of the scaffold itself is subtracted from the totalcalcium content of the lysate. The calcium concentration in the celllysates is quantified spectrophotometrically with cresolphthaleincomplexone (Sigma). Three minutes after the addition of reagents, theabsorbance of the samples is read at 575 nm using a microplate reader(Multiskan Spectrum, Thermo Electron Co., Vantaa, Finland). The calciumconcentration is calculated from a standard curve generated from aserial dilution of a calcium standard solution (Sigma).

The surface and cross-sectional morphologies of the scaffolds andcell-scaffold constructs can be examined using a SEM. The samples arewashed twice with PBS, prefixed in 1% (v/v) buffered glutaraldehyde for1 h, and fixed in 0.1% (v/v) buffered formaldehyde for 24 h. The fixedsamples are dehydrated in ascending grades of ethanol, dried, andmounted on aluminum stubs using double-sided carbon tape. The specimensare coated with gold using a Sputter Coater (Cressington 108,Cressington Scientific Instruments, Cranberry, Pa.) and examined withSEM at an acceleration voltage of 10 kV.

In addition to the culture of cell-scaffold constructs in vitro, cellscaffold constructs can be implanted into the subcutaneous space ofathymic mice (BALB/c-nu, 7 weeks old, female, SLC, Tokyo, Japan). Afterthe mice are anesthetized with an intramuscular administration ofketamine hydrochloride (50 mg/kg, Yuhan Co., Seoul, Korea) and xylazinehydrochloride (5 mg/kg, Bayer Korea Ltd., Seoul, Korea), small incisionsare made on the dorsal skins of six mice. Four pouches per animal aremade by blunt dissection in subcutaneous sites, and cell-seededscaffolds are immediately implanted into the pouches (n=4).Subsequently, the skin is closed with 5-0 Vicryl sutures (Ethicon,Lenneke Marelaan, Belgium). The mice are housed singly after surgery andreceived humane care in compliance with the Hanyang UniversityGuidelines for the care and use of laboratory animals. The implants areretrieved for analysis at five and eight weeks after implantation.

The mechanical properties of the scaffolds can be assessed usingcompressive and tensile mechanical tests.

To determine whether the scaffold fabrication process affects the extentof HA exposure at the scaffold surface, the exposed HA can be stainedwith a hydrophilic dye. The surface composition of the PLGA/HA compositescaffolds can be analyzed with XPS.

First, the GF process avoids the use of organic solvents. Residualorganic solvents remaining in scaffolds may damage transplanted cellsand surrounding tissues. Furthermore, exposure to organic solvents mayinactivate biologically active factors. Therefore, the GF process maycause less denaturation of the growth factors incorporated within thescaffolds.

Second, the GF method can efficiently expose bioceramics at the surfaceof the polymer/bioceramic composite scaffolds. Staining with ahydrophilic dye and XPS analysis would show that the GF method exposed asignificantly higher extent of HA at the scaffold surface than did theconventional SC/PL method. Therefore, a GF scaffold can increase thechances of osteogenic cells to make contact with the bioactive ceramics,which enhances osteoblast differentiation and growth.

FIGS. 1 and 15 are photographs of the bone graft in block form.

Example 2 Powder

PLGA/HA scaffolds were created using the process of Example 1 and wereground to a powder sized between about 250 nm and about 2 mm. Mostpreferable the powder is graded into two different sizes of particles,the smaller consisting particles between about 250 nm and about 750 nmand the larger being between about 750 nm and about 2 mm. The resultantmaterial can be used to pack surfaces around devices implanted in boneand packed in voids in bone to induce new bone growth. FIG. 14 is aphotograph of the bone graft in powder form.

Those of skill in the art are familiar with the use of bone graftmaterials. The present invention can be used in the same manner as anyother bone graft material. In a preferred delivery system, the bonegraft material is mixed with polyethylene glycol or another hydrogel toform a paste. The material can be premixed and sold in a syringe foreasy application or can be mixed at the point of use and delivered viaany convenient means. When provided in a dry state, any suitablebiocompatible fluid can be used to wet the material and create a pastefor administration to the patient. Examples of such biocompatible fluidwetting agents include, but are not limited to: dextrose, glucose,maltose or sodium chloride solutions, blood, serum, plateletconcentrate, bone marrow aspirate, and synovial fluid. A biologicalfluid can be used in the form obtained from the biological source, or itcan be processed by application of one or more desired usefultechniques, examples of which include, separation techniques, such asfiltration (macro-, micro-, or ultra-filtration); purificationtechniques, such as dialysis; concentration techniques; andsterilization techniques.

One of skill in the art will recognize that other biological componentsincluding but not limited to proteins, growth factors, cells, stemcells, osteoblasts or such other components that will promote bonegrowth or maturation of the bone graft.

Example 3 Molded Bone Grafts

The scaffold of Example 1 can be molded into a rigid implantable bonegraft material of desired shapes by creating an appropriate mold for usein the reaction vessel. Referring to FIG. 2-7, the mold 11 consists of abottom 40 and a right side 16 and a left side 17 which define a void 23.End caps 24 are placed at each end and a top cap 14 closes the mold. Theshape of the void 23 determines the final shape of the bone graft 10.The mold can be designed to produce a single implant or can be designedto produce multiple implants. In the figures the mold contains a dividerfor producing two implants. One of skill in the art will appreciate thatthe mold can be designed to create bone grafts having complex shapes aswell as simple geometric shapes. In the Figures, the molds are assembledusing screws 22 although clamps or shaping the mold so it is evenlysupported in the pressure vessel can also be used. When screws 22 areused, the top cap 14 and end caps 24 will have holes 18 drilled in themwhich correspond with threaded holes 18 in the right and left sides ofthe mold. Production molds 11 would be shaped to produce bone grafts 10which roughly correspond to the shapes of bones to be replaced to reducethe amount of additional shaping required for use. When molded in thisfashion, the need for further processing of the bone graft 10 is reducedor even possibly eliminated depending on the final use of the bone graft10.

Example 4 Teflon Sheaths

Because of the extreme pressures generated in the reaction vessel 25,the scaffolds 10 in Example 1 expand making extraction of the rigidscaffold 10 extremely difficult. Additionally, there were signs of metalcontamination of the bone graft material. To overcome these problemsTeflon sheaths 11 were designed to facilitate removal. Referring to FIG.2-7, the sheaths 11 can also be designed in a way to allow the scaffoldto be molded during production. The Teflon material expands underpressure in the mold and shrinks back to its original shape when broughtback to ambient temperature making it easy to remove the bone graft fromthe mold. Additionally the Teflon prevents contamination of the bonegraft with the metal of the mold. When molded in this fashion, the needfor further processing of the scaffolds is reduced or even possiblyeliminated depending on the final use of the scaffold.

Scaffolds were formed using the process of Example 1 and the molds ofExample 3 using a sheath 11 made of Teflon to line the reaction vessel25 to facilitate removal of the scaffolds after production. The sheath11 is machined or molded from Teflon and comprises a bottom 15, a rightside 20 and a left side 19 which define a void 21 for receiving thescaffold material, and a liner cap 13. The sheath fits inside the mold11.

In use, the sheath 12 would be placed inside the mold 11, filled withthe bone graft material and the mold closed before being placed in thereaction vessel.

Example 5 Water-Jet Cutting

PLGA melts at a relatively low temperature. When cut with conventionalcutting implements, the friction of the cutting surface causes the graftto melt destroying the pore structure. As such conventional toothedsaws, rotary cutting instruments and drills cannot be used to shape thematerial without destroying its properties. Water cooled tools do notreduce the heat fast enough at the point of contact with the tool toprevent melting. This has been a serious impediment to shaping PLGAbased materials. The problem has been overcome by using waterjets withor without an abrasive to perform the cutting.

The use of conventional cutting agents such as aluminum oxides orcarborundum is to be avoided in implants because they will contaminatethe implant. As a result, a non-conventional cutting agent had to beused. The inventor has discovered that HA is a suitable cutting agentand avoids the contamination issues of other cutting agents. The onlysize limitation on HA particles is the particle size limits of theequipment being used.

Waterjets are commonly used in the art and can be hand held or part ofmulti axis routers.

Scaffolds manufactured as in Examples 1 and 2 were cut into desiredshapes using a waterjet using hydroxyapatite in the water as a cuttingaid.

Example 6 Collagen Impregnation

Bone grafts manufactured and shaped according to Examples 1-3 wereimpregnated with collagen by forcing a collagen solution under pressureinto the scaffold. The collagen can be from any source which would notrender the bone graft unfit for its intended use. While common sourcesinclude pig, cow, and rat, human collagen is most preferred. Suitablecollagen can be purchased commercially from companies such as Sunmaxunder the trade name Porcogen. Such collagen is sold as a solution ofapproximately 0.01 N HCl at a concentration of about 3 mg/ml. The stocksolution is diluted by addition of approximately 9× phosphate buffer orcell culture media and the pH adjusted to approximately 7 by titrationwith 0.1N NaOH and/or 0.1N HCl. Encoll Corporation is another collagensupplier.

In another preferred embodiment the collagen is recombinant humancollagen. Recombinant collagen potentially avoids contamination orpurity issues which may exist with collagen processed from animal orhuman sources. A source of recombinant human collagens are FG-5016and/or FG-5030 from Fibrogen Corporation. FG-5-16 is a recombinant humancollagen where FG-5030 is a cross linked recombinant human collagen.

The collagen solutions should be between 0.1% and 10% with the mostpreferred range between approximately 2% and 5% collagen by weight inthe solution.

In one embodiment the collagen is infused in the bone grafts by placingthe bone graft in a centrifuge tube, adding the collagen solution andcentrifuging at the equivalent RCF of 5000 g or higher for 10 minutes.This produces an acceptable degree of collagen distribution in the bonegraft.

In another embodiment collagen is forced into the bone graft usingpressure. In this embodiment the bone grafts are removed from thereactor and their sheaths and inspected. Optionally, the grafts can befurther shaped at this time. The bone grafts were then placed into avacuum chamber and placed under a vacuum. A preferred device is acommercial lyophilizer used to freeze dry. The vacuum is maintaineduntil the moisture level is reduced to between 20% and 3% with 5% beinga preferred target.

Example 7 Bone Graft Modeling

The bone grafts 10 of the present invention can be engineered toprecisely replace damaged bone using CAD/CAM processes. In a first step,a patient needing bone replacement undergoes imaging to measure theregion needing replacement. The measurements can be made using anysuitable imaging technology from which three dimensional measurementscan be taken, but at the present time computer aided tomography (CAT)scanning is the preferred method. Cone Beam Computerized Tomography,(CBCT) is a good alternative. Magnetic Resonance Imaging (MRI) is apotential alternative.

The dimensions of the bone defect to be reconstructed obtained fromimaging software are exported into a solid file. (.stl, .igs or anyother solid file format used in the industry). This solid file is loadedinto the CAD software. The CAD software can then be used to design amold suitable for use within the reaction vessel and/or to design ascaffold which can be cut using computer aided manufacturing technology.In some instances a mold may be sufficient to produce a final shape. Inother situations, further shaping of the materials may be required. Thescaffolds (molded or otherwise) can be shaped on a multi axis router.Because the polymers of the scaffolds can melt if exposed to a source ofheat, the cutting heads need to be cooled. In a most preferredembodiment the cutting is performed using a waterjet, preferably withthe use of hydroxyapatite particles as a cutting abrasive.

The resulting bone grafts 10 would be processes for storage and shippingas is customary in the art. Such processing steps include, but are notlimited to, freeze drying, packaging and sterilization.

The resultant shaped bone graft 10 would then be implanted by a surgeoninto a patient following removal of the area of damaged bone. Thedamaged bone can be removed to fit the bone graft or the bone graft canbe further shaped by the surgeon using common cutting tools or abrasivesto fit the damaged area.

Example 8 Spinal Fusion

Bone grafts of the present invention may be used for spinal fusion, aprocess in which two or more vertebrae are connected together.Traditional, surgeons place screws in each of the vertebrae to be fusedtogether and connect them with plates or rods to prevent movement. Boneis preferable grafted in between the vertebrae to facilitate fasterfusion. The grafts of the present invention can be shaped to morereadily fit between vertebrae. The present invention facilitates fasterfusion of the vertebral segments than using harvested bone.

The pores of the present invention also offer the ability to seed thebone graft with stem cells or biological agents. In one instance, thescaffold could be seed with cells to regrow cartilage and or ligaments.While such an implant would initially be rigid and unyielding, becausethe PLGA is slowly resorbed eventually the implant would gain thecushioning and flexibility of a normal disc.

Example 9 Bone Graft Kits

Because exact shape matching of a bone graft 10 to the region of bone tobe replaced may be difficult prior to surgery, a kit can be createdhaving various sizes of bone grafts. The bone graft selection can bemade by the surgeon during the procedure. Such kits could include aplurality of scaffold shapes and sizes, suitable bone cements, as wellas surgical tools, screws, shaping aids etc. which may be required toshape the scaffold before implanting, remove bone or otherwise berequired during the surgery.

The description of the teachings is merely exemplary in nature and,thus, variations that do not depart from the gist of the teachings areintended to be within the scope of the teachings. Such variations arenot to be regarded as a departure from the spirit and scope of theteachings.

The description of the teachings is merely exemplary in nature and,thus, variations that do not depart from the gist of the teachings areintended to be within the scope of the teachings. Such variations arenot to be regarded as a departure from the spirit and scope of theteachings.

REFERENCES

-   1. de Boer H H. The history of bone grafts. Clin Orthop Relat Res    1988; 226:292-8.-   2. Vacanti C A, Kim W, Upton J, Vacanti M P, Mooney D, Schloo B, et    al. Tissue-engineered growth of bone and cartilage. Transplant Proc    1993; 25:1019-21.-   3. Bonfiglio M, Jeter W S. Immunological responses to bone. Clin    Orthop Relat Res 1972; 87:19-27.-   4. Coombes A G, Meikle M C. Resorbable synthetic polymers as    replacements for bone graft. Clin Mater 1994; 17:35-67.-   5. Rizzi S C, Heath D J, Coombes A G, Bock N, Textor M, Downes S.    Biodegradable polymer/hydroxyapatite composites: surface analysis    and initial attachment of human osteoblasts. J Biomed Mater Res    2001; 55:475-86.-   6. Laurencin C T, Attawia M, Borden M D. Advancements in tissue    engineered bone substitutes. Curr Opin Orthop 1999; 10:445-51.-   7. Ambrosio A M, Sahota J S, Khan Y, Laurencin C T. A novel    amorphous calcium phosphate polymer ceramic for bone repair: I.    Synthesis and characterization. J Biomed Mater Res 2001; 58:    295-301.-   8. Marra K G, Szem J W, Kumta P N, DiMilla P A, Weiss L E. In vitro    analysis of biodegradable polymer blend/hydroxyapatite composites    for bone tissue engineering. J Biomed Mater Res 1999; 47:324-35.-   9. Wang M. Developing bioactive composite materials for tissue    replacement. Biomaterials 2003; 24:2133-51.-   10. Van Landuyt P, Li F, Keustermans J P, Streydio J M, Delannay F,    Munting E. The influence of high sintering temperatures on the    mechanical properties of hydroxyapatite. J Mater Sci Mater Med 1995;    6:8-13.-   11. Khan Y M, Katti D S, Laurencin C T. Novel polymer-synthesized    ceramic composite-based system for bone repair: An in vitro    evaluation. J Biomed Mater Res A 2004; 69:728-37.-   12. Kikuchi M, Cho S-B, Suetsugu Y, Tanaka J. In vitro tests and in    vivo tests developed TCP/CPLA composites. Bioceramics 1997; 10:    407-10.-   13. Reis R L, Cunha A M, Fernandes M H, Correia R N. Bioinert and    biodegradable polymeric matrix composites filled with bioactive    SiO2-3CaO    P2O5-MgO glasses and glass-ceramics. Bioceramics 1997; 10:415-8.-   14. Piattelli A, Franco M, Ferronato G, Santello M T, Martinetti R,    Scarano A. Resorption of composite polymer-hydroxyapatite membranes:    a time-course study in rabbit. Biomaterials 1997; 18: 629-33.-   15. Lu L, Currier B L, Yaszemski M J. Synthetic bone substitutes.    Curr Opin Orthop 2000; 11:383-90.-   16. Peter S J, Lu L, Kim D J, Mikos A G. Marrow stromal osteoblast    function on a poly(propylene fumarate)/beta-tricalcium phosphate    biodegradable orthopaedic composite. Biomaterials 2000; 21:1207-13.-   17. Wei G, Ma P X. Structure and properties of    nano-hydroxyapatite/polymer composite scaffolds for bone tissue    engineering. Biomaterials 2004; 25:4749-57.-   18. Guan L, Davies J E. Preparation and characterization of a highly    macroporous biodegradable composite tissue engineering scaffold. J    Biomed Mater Res A 2004; 71:480-7.-   19. Zhang R, Ma P X. Poly(alpha-hydroxyl acids)/hydroxyapatite    porous composites for bone-tissue engineering. I. Preparation and    morphology. J Biomed Mater Res 1999; 44:446-55.-   20. Lee S H, Kim B S, Kim S H, Kang S W, Kim Y H. Thermally produced    biodegradable scaffolds for cartilage tissue engineering. Macromol    Biosci 2004; 4:802-10.-   21. Yang S, Leong K F, Du Z, Chua C K. The design of scaffolds for    use in tissue engineering. Part I. Traditional factors. Tissue Eng    2001; 7:679-89.-   22. Jung Y, Kim S S, Kim Y H, Kim S H, Kim B S, Kim S, et al. A    poly(lactic acid)/calcium metaphosphate composite for bone tissue    engineering. Biomaterials 2005; 26:6314-22.-   23. Jung Y, Kim S H, Kim S S, You H J, Kim B S, Kim S, et al. Tissue    engineered bone formation with polymer/ceramic composites by    press-and-baking method. Key Eng Mater 2005; 288:79-82.-   24. Harris L D, Kim B S, Mooney D J. Open pore biodegradable    matrices formed with gas foaming. J Biomed Mater Res 1998; 42:    396-402.-   25. Cho S W, Kim I K, Lim S H, Kim D I, Kang S W, Kim S H, et al.    Smooth muscle-like tissues engineered with bone marrow stromal    cells. Biomaterials 2004; 25:2979-86.-   26. Cho S W, Kim S S, Rhie J W, Cho H M, Choi C Y, Kim B S.    Engineering of volume-stable adipose tissues. Biomaterials 2005; 26:    3577-85.-   27. Kim B S, Jeong S I, Cho S W, Nikolovski J, Mooney D J, Lee S H,    et al. Tissue engineering of smooth muscle under a mechanically    dynamic condition. J Microbiol Biotech 2003; 13:841-5.-   28. Whitson S W, Whitson M A, Bowers Jr. D E, Falk M C. Factors    influencing synthesis and mineralization of bone matrix from fetal    bovine bone cells grown in vitro. J Bone Miner Res 1992; 7:727-41.-   29. Ekholm M, Hietanen J, Tulamo R M, Muhonen J, Lindqvist C,    Kellomaki M, et al. Tissue reactions of subcutaneously implanted    mixture of epsilon-caprolactone-lactide copolymer and tricalcium    phosphate. An electron microscopic evaluation in sheep. J Mater Sci    Mater Med 2003; 14:913-8.-   30. Jaiswal N, Haynesworth S E, Caplan A I, Bruder S P. Osteogenic    differentiation of purified, culture-expanded human mesenchymal stem    cells in vitro. J Cell Biochem 1997; 64:295-312.-   31. Lewandrowski K U, Bondre S P, Wise D L, Trantolo D J. Enhanced    bioactivity of a polypropylene fumarate) bone graft substitute by    augmentation with nano-hydroxyapatite. Biomed Mater Eng 2003;    13:115-24.-   32. Ginebra M P, Driessens F C, Planell J A. Effect of the particle    size on the micro and nanostructural features of a calcium phosphate    cement: a kinetic analysis. Biomaterials 2004; 25: 3453-62.-   33. Burg K J L, Porter S, Kellam J F. Biomaterial Developments for    bone tissue engineering. Biomaterials 2000; 21:2347-2359.-   34. Akoa M, Aoki H, Kato K. Mechanical properties of sintered    hydroxyapatite for prosthetic applications. J Mater Sci 1981;    16:809-812.-   35. Anselme K. Osteoblast adhesion on biomaterials. Biomaterials    2000; 21:667-681.-   36. Howe A K, Aplin A E, Juliano R L. Anchorage-dependent ERK    signaling-mechanisms and consequences. Curr Opin Genet Dev 2002;    12:30-35.-   37. Bigi A, Boanini E, Panzavolta S, Roveri N, Rubini K. Bone like    apatite growth on hydroxyapatite-gelatin sponges from simulated body    fluid. J Biomed Mater Res 2002; 59:709-715.-   38. Stupp S I, Ciegler G W. Organoapatites: Materials for artificial    bone. I. Synthesis and microstructure. J Biomed Mater Res 1992;    26:169-183.-   39. Vandiver J, Dean D, Patel N, Bonfield W, Ortiz C. Nanoscale    variation in surface charge of synthetic hydroxyapatite detected by    chemically and spatially specific high-resolution force    spectroscopy. Biomaterials 2005; 26:271-283.-   40. Lu H H, El-Amin S F, Scott K D, Laurencin C T.    Three-dimensional, bioactive, biodegradable, polymer-bioactive glass    composite scaffolds with improved mechanical properties support    collagen synthesis and mineralization of human osteoblast-like cells    in vitro. J Biomed Mater Res A 2003; 64: 465-474.-   41. Li H, Chang J. Preparation and characterization of bioactive and    biodegradable wollastonite/poly(D,L-lactic acid) composite    scaffolds. J Mater Sci Mater Med 2004; 15:1089-1095.

1-56. (canceled)
 57. A bone graft biomaterial comprised of a scaffold consisting of a biocompatible polymer and a bioceramic composite in a ratio of about 1:2 to about 2:1 wherein the bioceramic particles are less than 1000 nm in diameter and wherein the scaffolds contain interconnected pores made via a gas foaming or a particulate leaching process, wherein the bioceramic particles are exposed on the surface of the biomaterial and wherein the biomaterial is impregnated with collagen after the pores are formed by forcing the collagen into the pores by placing the scaffold into a collagen solution and the solution is placed under a vacuum until the temperature is just above the freezing point of the collagen solution, the solution allowed to warm before being subjected to repeated cycles of vacuum and warming.
 58. The bone graft material of claim 57 wherein the biocompatible polymer is selected from the group comprising Poly lactic acid (PLA), poly glycolic acid (PGA), Poly lactic co-glycolic acid (PLGA), and copolymers with polyethylene glycol (PEG); polyanhydrides, poly(ortho)esters, polyurethanes, poly(butyric acid), poly(valeric acid), poly(lactide-co-caprolactone) and trimethylene carbonate and combinations and co-polymers thereof.
 59. The bone graft material of claim of claim 57 wherein the bioceramic is selected from the group consisting of hydroxyapatite, tricalcium phosphate, bioglass, calcium phosphate or bone or a combination thereof.
 60. The bone graft material of claim 57 wherein the bioceramic particles are from 100 nm to 1000 nm in diameter.
 61. The bone graft material of claim 60 wherein the bioceramic particles are from 100 nm to 400 nm in diameter.
 62. The bone graft material of claim 57 wherein the biocompatible polymer is poly lactic acid, poly glycolic acid, lactide or caprolactone or copolymers or combinations thereof.
 63. The bone graft material of claim 62 comprising two polymers in a ratio from 2:1 to 1:2.
 64. The bone graft material of claim 62 where the polymers are lactide and caprolactone in a ratio of about 75:25 by weight.
 65. A bone graft biomaterial of claim 57 wherein the bioceramic is TCP coated hydroxyapatite.
 66. The bone graft material of claim 57 wherein the ratio of biocompatible polymer to bioceramic ranges from about 80:20 to about 50:50.
 67. The bone graft material of claim 64 wherein the ratio of polymer to bioceramic is about 60:40
 68. The bone graft material of claim 59 where in the bioceramic particles are a mixture of hydroxy apatite and TCP in a ratio of about 40:60.
 69. A bone graft material comprising lactide and caprolactone polymers in a ratio from about 1:2 to about 2:1 which are in a ratio of from about 80:20 to about 50:50 with a bioceramic having a diameter from about 150 microns to about 300 microns comprising hydroxyapatite, TCP, hydroxyapatite TCP or a mixture thereof wherein the scaffold is formed via gas foaming.
 70. The bone graft material of claim 69 where in the ratio of lactide to caprolactone is 75:25.
 71. The bone graft material of claim 69 wherein the ratio of lactide and caprolactone polymers to bioceramic is about 60:40.
 72. The bone graft material of claim 57 wherein the material is compressed using an inert gas at pressures from about 400 psi to about 3000 psi for a period of about 1 to about 8 hours, before reducing pressure to ambient with a period of about 10 minutes to about 30 minutes.
 73. The bone graft material of claim 72 wherein the material is compressed using an inert gas at pressures from about 1500 to about 2500 PSI for a period of about 1-5 hours before reducing pressure to ambient with a period of about 10 minutes to about 30 minutes.
 74. A method of fabricating a polymer and hydroxyapatite biomaterial scaffold for use in replacing bone comprising: a) introducing a biocompatible polymer selected from selected from the group comprising Poly lactic acid (PLA), poly glycolic acid (PGA), Poly lactic co-glycolic acid (PLGA), and copolymers with polyethylene glycol (PEG); polyanhydrides, poly(ortho)esters, polyurethanes, poly(butyric acid), poly(valeric acid), poly(lactide-co-caprolactone) and trimethylene carbonate and combinations and co-polymers thereof and a bioceramic selected from of hydroxyapatite, tricalcium phosphate, bioglass, calcium phosphate or bone or a combination thereof into a mold in a pressure reactor; b) pressurizing the reactor with an inert gas to form a scaffold; c) reducing the pressures to induce pore forming bubbles in the scaffold; and d) introducing the scaffold to a collagen solution; e) driving the collagen solution into the scaffold using pressure or a vacuum; and f) optionally shaping the scaffold using cutting tools.
 75. The method of claim 74 wherein the mold further comprises a Teflon sheath.
 76. The method of claim 75 wherein the scaffold is shaped with a water jet.
 77. The method of claim 74 having the following steps prior to step a: i) imaging a region in a patient requiring bone replacement to obtain the dimensions of bone requiring replacement ii) using the dimensions of the bone requiring replacement to determine the dimensions of the required bone graft iii) fabricating a mold for a pressure reactor using the calculated bone replacement dimensions. 